MR angiography using steady-state transport-induced adiabatic fast passage

ABSTRACT

A subtractive time of flight technique for MR angiography and quantitative blood flow measurement. Proton spins of water in the arterial supply to a tissue or organ are inverted in a steady-state fashion by applying constant amplitude off-resonance radio frequency pulses in the presence of a constant linear magnetic field gradient to effect adiabatic fast passage transport-induced inversion of spins which move in the direction of the gradient. An angiogram is formed by subtracting an image acquired with the arterial inversion pulse from a control image acquired with no arterial inversion. By inverting the spins in a steady-state manner, no cardiac gating is necessary for imaging organs. However, cardiac gating is desirable when imaging the heart so that spins of blood passing through the coronary arteries can be inverted during systole, when most of the blood is in the left ventricle, and imaged at end diastole, when most of the blood is in the coronary arteries. A coronary angiogram is then formed by subtracting images acquired with and without the inversion pulse. Also, by applying several inverting and imaging pulses during a cardiac cycle in accordance with the technique of the invention, a characteristic banding pattern may be formed in the fluid whereby each band corresponds to a population of spins that experienced inversion due to a single RF pulse. Since the width of the inversion band is proportional to the duration of the RF pulse and the velocity of the spin, measurement of the thickness of the inverted and uninverted bands allows for calculation of flow velocity. By gating such a pulse sequence to the cardiac cycle, time resolved in vivo velocity measurements may be made.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to a technique for MR angiography andquantitative blood flow measurement, and more particularly, totechniques for MR angiography using steadystate transport-inducedadiabatic fast passage and for quantitative MR flow measurement usingpulsed adiabatic inversion.

2. Description of the Prior Art

The measurement of tissue perfusion, i.e., the flow of fluid in tissue,and of quantitative blood flow in arteries is important for thefunctional assessment of organs in vivo. Although the terms perfusionand flow are sometimes used interchangeably, perfusion as used herein isa quantifiable measurement of capillary blood flow which is generallymeasured indirectly in humans, while flow refers to the quantitativevolume of blood passing through an artery in a given period of time.Angiograms, on the other hand, provide a qualitative view of arteries,tissues and the like which is useful for diagnostic purposes. Numeroustechniques have been developed in the prior art for providing angiogramsas well as measurements of perfusion and blood flow using magneticresonance imaging.

For example, in U.S. Pat. application Ser. No. 746,771, filed Aug. 16,1991, and in an article entitled "Magnetic Resonance Imaging ofPerfusion Using Spin Inversion of Arterial Water", Proc. Natl. Acad.Sci., Vol. 89, pp. 212--216 (1992), one of the present inventorsdisclosed a method for measuring perfusion by labeling proton spins ofinflowing water in the arterial blood using magnetic resonance.Continuous saturation or continuous inversion using an adiabaticexcitation was then performed proximal to the tissue or organ ofinterest. In particular, perfusion was measured by labeling atoms in theblood at a base point, generating a steady state in the tissue or organby continuing to label atoms until the effect caused by labeled atomsperfusing into the tissue or organ reaches a steady state, generatingimaging information for the tissue or organ, and processing the imaginginformation to determine perfusion. In a particular embodiment of thatinvention, by continuously applying a radio-frequency (RF) field, spinsassociated with the atoms were inverted continuously by adiabatic fastpassage. This technique was disclosed for determining perfusion in thebrain as well as other tissues or organs having well defined arterialsupplies such as the kidney, the liver and the heart.

Spin inversion was preferred for perturbing the magnetization of thearterial water in accordance with the technique of the above-identifiedapplication because it maximized the observed effect. In order to invertthe arterial spins continuously as described therein, an RF signal wasapplied in the presence of a magnetic field gradient in the direction ofthe flow so that the movement of the spins through the magnetic fieldgradient leads to a change in the magnetic field through resonance. Asdescribed therein, this approach has lead to greatly improved images oftissue perfusion in the brain and other organs.

However, since the contribution from the arterial intravascular volumemust be eliminated in order to get accurate measurements of tissueperfusion, the technique described in the aforementioned patentapplication is not suitable for quantitatively measuring blood flow orfor acquiring an MR angiogram. On the contrary, the technique describedin the aforementioned patent application requires symmetrical spoilergradient pulses to be used in the imaging sequence around the 180° pulseto eliminate arterial intravascular signals in both the proximal and thecontrol saturation images. This minimizes the affects on intravascularspins of proximal saturation, which otherwise causes the differenceimage to include a contribution from the arterial intravascular volumein addition to a contribution caused by the exchange of labeled vascularwater with tissue water during the perfusion measurement. In otherwords, spoiler gradients have been used when obtaining images of tissueperfusion in order to eliminate signal contribution from the movingspins. For this reason, while providing an excellent technique formeasuring perfusion, the technique described in the aforementionedpatent application cannot be used for MR angiography and quantitativeblood flow measurement.

Recently, several subtractive time-of-flight (TOF) magnetic resonanceangiography (MRA) techniques have been described in the art. Forexample, such techniques are described by Dixon et al. in an articleentitled "Projection Angiograms of Blood Labeled by Adiabatic FastPassage", Magnetic Resonance in Medicine, Vol. 3, pp. 454-462 (1986); byNishimura et al. in articles entitled "MR Angiography By SelectiveInversion Recovery", Magnetic Resonance in Medicine, Vol. 4, pp. 193-202(1987), "Considerations of Magnetic Resonance Angiography By SelectiveInversion Recovery", Magnetic Resonance in Medicine, Vol. 7, pp. 472-484(1988), and "Magnetic Resonance Angiography By Selective InversionRecovery Using A Compact Gradient Echo Sequence", Magnetic Resonance inMedicine, Vol. 8, pp 96-103 (1988); by Sardashti et al. in an articleentitled "Spin-Labeling Angiography of the Carotids By Presaturation andSimplified Adiabatic Inversion", Magnetic Resonance in Medicine, Vol.15, pp. 192-200 (1990); and by Wang et al. in an article entitled "FastAngiography using Selective Inversion Recovery", Magnetic Resonance inMedicine, Vol. 23, pp. 109-121 (1992). The techniques described in thosearticles complement conventional phase-contrast techniques such as thosedescribed by Dumoulin et al. in an article entitled "Magnetic ResonanceAngiography", Radiology, Vol. 161, pp. 717-720 (1986); by Wedeen et al.in an article entitled "Projective MRI Angiography and QuantitativeFlow-Volume Densitometry", Magnetic Resonance in Medicine, Vol. 3, pp.226-241 (1986); and by Nayler et al. in an article entitled "Blood FlowImaging by Cine Magnetic Resonance", Journal of Computer AssistedTomography, Vol. 10, pp. 715-722 (1986), and conventional time-of-flightMRA techniques such as those described by Gullberg et al. in an articleentitled "MR Vascular Imaging With a Fast Gradient Refocusing PulseSequence and Reformatted Images From Transaxial Sections", Radiology,Vol. 165, pp. 241-246 (1987).

Dixon et al. describe an approach to MR angiography in which a constantRF pulse is transmitted through a separate surface coil in the presenceof a constant magnetic gradient so that moving spins undergo adiabaticfast passage inversion. By turning the inversion RF on and off duringalternate gated acquisitions, Dixon et al. generated images with andwithout labeled blood. An angiogram was then formed by simplesubtraction for visualizing, for example, the carotid bifurcation.

In accordance with the Dixon et al. technique, all the blood passing theRF surface coil while the RF signal is being applied is labeled and thenallowed to flow for a period of time so that it can enter the imagingvolume. The longest possible labeling time is used which is consistentwith the number of cardiac cycles chosen for TR. It is thus important inusing the Dixon et al. technique that the RF signal be turned on and offslowly to prevent saturation of stationary tissue near the RF surfacecoil. It is also important that little power be used so as to preventheating of the tissue of the patient in the vicinity of the RF surfacecoil. Moreover, in accordance with the technique of Dixon et al., sincethe phase encodings are interleaved and the labeled blood must be giventime to propagate into the imaging volume before acquiring the image,only one phase encoding per cardiac cycle is possible. Furthermore,since this process is synchronized to the cardiac cycle by cardiacgating so as to phase encode every cardiac cycle with 90° and 180°pulses, the appropriate propagation time delay must also be selected toallow the labeled column of blood enough time to fill the arteries thatare being imaged. Accordingly, Dixon et al. must estimate how far thespins travel in a selected period of time in order to maintainsynchronization. An improved approach is desired which does not requiresynchronization to the cardiac cycle and which allows for fast scanimaging.

Another approach to MRA, referred to as selective inversion recovery(SIR) (as described in the aforementioned articles to Nishimura et al.,Sardashti et al. and Wang et al.), uses spatially selective inversionpulses to label a column of arterial blood which then flows into anorgan in the imaging region. Selective 180° excitation inverts differentregions between measurements to isolate arterial and/or venous blood sothat high-resolution carotid artery angiograms and the like may beobtained. A 90° presaturation pulse may also be applied to suppress thebackground intensity from static tissue in the angiogram, which isformed by subtracting an image acquired with inversion from an imageacquired without inversion. As an inversion technique, SIR of thearterial inflow to an organ has been found to offer excellentangiographic contrast and penetration due to the fact that blood isinverted, not just saturated. As a subtractive technique, SIR alsoprovides excellent background suppression, similar to that of phasecontrast techniques.

However, there are also some problems associated with SIR techniqueswhich use spatially selective inversion pulses to label blood flow. Ingeneral, these techniques must be gated to the cardiac cycle so thatcomplete labeling of the inflowing blood occurs. In addition, theinversion pulses used to label blood flow should have precisely definedpassbands in order for complete inversion to occur. This typicallyinvolves the use of complicated RF pulse shapes. Another drawback of SIRis the fact that spins in the more distal parts of the labeled column ofblood undergo T1-relaxation as they traverse the slab thickness. Thespins therefore lose part of the label before entering the imagingvolume, leading to a loss of angiographic contrast. An improvedangiographic technique is desired which overcomes such problems.

Various time-of-flight (TOF) methods have also been described in theprior art for quantitative NMR flow measurement. For example, Saloner etal. disclose in an article entitled "Flow Velocity Quantitation UsingInversion Tagging", Magnetic Resonance in Medicine, Vol. 16, pp. 269-279(1990) and Edelman et al. disclose in an article entitled"Quantification of Blood Flow With Dynamic MR Imaging and PresaturationBolus Tracking", Radiology, Vol. 171, pp. 551-556 (1989) methodsemploying spatially localized saturation or inversion RF pulses forlabeling moving spins. In these approaches, a column of fluid issubjected to a series of spatially-selective RF pulses in a steady-stateimaging experiment. This results in a banding pattern in the fluid fromwhich the velocity may be determined. However, these methods are subjectto errors arising from motion of the spins during application of thespatially selective RF pulses. In addition, the accuracy of thesemethods also depends on the characteristics of the spatially-selectiveinversion pulses used to tag the fluid. In general, optimized RF pulsesmust be used in order for accurate velocity measurements to be made.Unfortunately, due to the fact that each group of "tagged" spinsexperiences multiple RF pulses during transit through the inverted slab,contrast between bands decreases as the flow moves downstream. Thiseffect makes it more difficult to observe the evolution of the bandingpattern as the fluid flows distal to the inversion plane.

Lee et al. describe another flow measurement technique in an articleentitled "Spatially Resolved Flow Velocity Measurements and ProjectionAngiography by Adiabatic Passage", Magnetic Resonance Imaging, Vol. 9,pp. 115-127 (1991), which provides direct assessment of in-plane andoblique directional flow velocities and visualization of flow velocityprofiles as well as flow angiography based on the time-of-flighttechnique. In particular, Lee et al. generate a band in the liquid whichmay be used for measuring flow velocity by applying a fast adiabaticpassage pulse after a cardiac gated sequence and then applying a spinecho sequence for acquiring the image. A 180° RF pulse is also appliedfor stationary spin suppression prior to the adiabatic fast passagepulse and a 90° RF pulse is applied prior to phase encoding.Unfortunately, Lee et al. do not sequence fast enough to see pluralbands in a single cardiac cycle, and since the 90° and 180° gradientpulses are used, fast scanning techniques cannot be used. Hence, thesystem of Lee et al. does not provide optimized time resolution of theimages taken during the cardiac cycle. An improved MR flow measurementtechnique is desired which will allow for fast scanning of the slab soas to improve time resolution.

The present invention has been designed to overcome the aforementionedlimitations in the prior art by providing a comprehensive system for MRangiography and MR blood flow measurement for different regions of thebody. As will be clear from the following description, the presentinventors have found that improved MR images of arteries, veins andtissues may be produced using a steady-state adiabatic inversiontechnique and that blood flow may be measured by adiabatically invertingspins in the direction of the magnetic gradient. The present inventionis thus believed to meet long felt needs in the prior art in that itprovides for the first time a methodology which allows for MRangiography as well as quantitative blood flow and tissue perfusionmeasurement of any tissue or organ having a well defined arterialsupply.

SUMMARY OF THE INVENTION

The above-mentioned long-felt needs of the prior art have been met inaccordance with the present invention by providing a subtractivetime-of-flight technique which is related to the pulse inversionapproach described by Dixon et al. except that in accordance with thetechnique of the invention the blood supply to an organ is continuouslyinverted as it flows into the imaging region. This is accomplished byapplying a off-resonance RF pulse in the presence of a constant magneticfield gradient. Since the RF pulse is applied in a steady-state imagingsequence, cardiac gating is not necessary. Also, since a single coil isused for application of the inversion and observation pulses inaccordance with the invention, inversion of the arterial supply mayoccur very close to the imaging volume, thereby minimizing transit timeeffects. This adiabatic spin inversion angiography technique of theinvention has been used to perform intracranial MRA and coronary MRA at1.5 Tesla, although those skilled in the art will appreciate that thetechnique of the invention may be applied to other tissues and organs aswell.

When imaging the vasculature of an organ in accordance with thetechniques of the invention, a single coil can be used to invert bloodat an arbitrary location in the imaging volume using the principle oftransport-induced adiabatic fast passage inversion. However, due to thecomplex geometry of the heart, a "localizer" image must first be takenwhich allows the MR operator to choose an imaging plane in the lateralanterior descending artery or an imaging plane which transects theaortic valve of the heart on its longitudinal axis across the entirecurve of the aorta and to establish an inversion plane at a pointinferior to the aortic valve so that the resulting angiogram will bemeaningful. Then, by applying the inversion pulse in the inversion planeduring ejection of blood by the left ventricle (systole), all bloodejected into the coronary arteries may be labeled. The imaging sequenceis then applied to the imaging plane in late diastole in order toobserve the blood in the coronary arteries. In this manner, the presentinvention makes possible for the first time coronary angiography as wellas coronary blood flow measurement using magnetic resonance.

A preferred embodiment of the method of generating an angiographic imageof a body portion of a patient using magnetic resonance in accordancewith the invention preferably comprises the steps of:

applying to the body portion an external substantially uniform magneticfield so as to align predetermined nuclei of the body portion with aconstant magnetic field gradient of the magnetic field, thepredetermined nuclei having a particular resonance frequency;

defining an inversion plane in the body portion which intersects a bloodsupply in a direction transverse to the magnetic field gradient byapplying via an RF coil disposed adjacent to the body portion asubstantially constant amplitude tagging radio-frequency (RF) pulsehaving a frequency different from the particular resonance frequency,the tagging RF pulse being applied in the presence of the constantmagnetic field gradient so as to invert spins of those predeterminednuclei which have a component of velocity in a direction of the magneticfield gradient as the predetermined nuclei pass through the inversionplane;

phase encoding the predetermined nuclei as they pass through theinversion plane during the tagging RF pulse so as to establish a steadystate inversion condition for the predetermined nuclei;

applying via the RF coil an observation RF pulse for acquiring a taggedimage of the body portion in a region near the inversion plane; and

generating an angiogram of the body portion in the region by effectivelysubtracting the tagged image of the region from an untagged image of theregion.

Also, by using a simple time-of-flight technique for

BRIEF DESCRIPTION OF THE DRAWINGS

The above and other objects and advantages of the invention will becomemore apparent and more readily appreciated from the following detaileddescription of the presently preferred exemplary embodiments of theinvention taken in conjunction with the accompanying drawings, of which:

FIGS. 1(a), 1(b) and 1(c) together illustrate the effect of a constantRF pulse on a spin moving through a gradient at constant velocity,whereby as a spin moves through the gradient, G_(i), it undergoes afrequency sweep as its instantaneous frequency changes linearly withtime.

FIG. 2 illustrates the simulated inversion efficiency, α, versus theamplitude, B_(i), of the inversion pulse for three different gradientstrengths.

FIGS. 3(a) and 3(b) together illustrate the effect of magnetizationtransfer on the technique of the invention where it is assumed that thecontrol image is acquired by changing the sign of an inversion gradientin the slabselection direction.

FIGS. 4(a), 4(b), 4(c) and 4(d) together illustrate a two-dimensionalaxial gradient-echo pulse sequence used to generate images in accordancewith the embodiment of the invention described with respect to Example1.

FIG. 5 illustrates the location of the image slab and inversion planefor use in producing an intracranial angiogram in accordance with thetechnique of the invention.

FIGS. 6(a) and 6(b) illustrate respective two-dimensional projectiveangiograms generated by changing the sign of the inversion gradient(FIG. 6(a)) and by changing the sign of the offset frequency of theinversion pulse (FIG. 6(b)).

FIGS. 7(a) and 7(b) together illustrate the effect of inversion gradientstrength on contrast, where G_(i) =0.25 gauss/cm (FIG. 7(a)) and G_(i)=1.00 gauss/cm (FIG. 7(b)).

FIGS. 8(a) and 8(b) together illustrate the effect of inversion pulseamplitude on contrast, where B_(i) =40 Hz (FIG. 8(a)) and B_(i) =120 Hz(FIG. 8(b)).

FIGS. 9(a), 9(b) and 9(c) together illustrate the time courses ofventricular ejection and coronary blood flow during the cardiac cycle.

FIGS. 10(a), 10(b), 10(c) 10(d) and 10(e) together illustrate a pulsesequence used to form coronary angiograms using transport-inducedadiabatic fast passage whereby a constant-amplitude off-resonance RFpulse in the presence of a constant gradient in the frequency-encodingdirection proceeds a standard two-dimensional gradient-echo imagingsequence.

FIG. 11 illustrates a coronary angiogram produced in accordance with theinvention whereby the bright horizontal band in the image indicates thegeometry of the inversion plane.

FIGS. 12(a), 12(b), 12(c) and 12(d) together illustrate the effect of aconstant RF pulse on a spin moving through a gradient at constantvelocity, whereby as the spin moves through the gradient, G, itundergoes a frequency sweep. FIG. 12 illustrates how the location of theinversion plane depends on the inversion gradient G_(i) and the offsetfrequency ω_(i).

FIG. 13(a) illustrates a typical banding pattern established in aflowing liquid by a series of inversion pulses in accordance with theinvention, where T_(i) is the duration of the inversion pulse, T_(ed) isthe time between the end of the inversion pulse and the center of theecho, and TR is the repetition time.

FIG. 13(b) illustrates the RF pulse sequence for forming the bands ofFIG. 13(a).

FIGS. 14(a), 14(b), 14(c) and 14(d) illustrate a two-dimensional coronalgradient-echo pulse sequence used to make ungated flow measurements inaccordance with the invention, whereby the inversion pulse is defined byits duration, T_(i), its strength, B_(i), and its frequency offset,ω_(i).

FIGS. 15(a), 15(b) and 15(c) illustrate the magnitude of gradient-echoimages of water flowing through 1/2 inch tubing, where B_(i) =20 Hz, 40Hz, and 80 Hz, respectively. As illustrated, bright inversion bands areformed at the higher RF frequencies.

FIG. 16 illustrates the measured (solid) and predicted (dashed) valuesfor velocity versus position across the lumen for steady flow in a 1/2inch tube phantom as illustrated in FIG. 15.

FIGS. 17(a), 17(b), 17(c), 17(d) and 17(e) together illustrate atwo-dimensional gated pulse sequence used to make in vivo flowmeasurements in accordance with the invention, whereby the inversionflow-tagging pulse precedes each phase encoding.

FIGS. 18(a) and 18(b) illustrate gated magnitude gradient-echo images ofthe internal carotid artery of a human patient using the pulse sequenceof FIG. 17 for the systolic portion of the cardiac cycle (FIG. 18(a))and the diastolic portion of the cardiac cycle (FIG. 18(b)).

FIG. 19 illustrates flow velocity as a function of trigger delay in theright internal carotid artery by measuring the thickness of the firstinflow band in a series of images gated to the cardiac cycle.

FIG. 20 illustrates pulse simulations of the effect of increasinggradient strength on adiabatic flow inversion in accordance with theinvention.

DETAILED DESCRIPTION OF THE PRESENTLY PREFERRED EXEMPLARY EMBODIMENTS

The MR angiography and MR flow measurement techniques in accordance withthe presently preferred exemplary embodiments of the invention will bedescribed below with reference to FIGS. 1-20. The MR angiographytechnique of the invention using steady-state transport-inducedadiabatic fast passage will first be described in connection withExample 1 (intracranial angiography) and FIGS. 1-8. Then, coronaryangiography using transport-induced adiabatic fast passage in accordancewith the invention will be described in connection with Example 2(coronary angiography) and FIGS. 9-11. Finally, a technique forquantitative NMR flow measurement using pulsed adiabatic inversion inaccordance with the invention will be described with respect to Example3 for phantom studies (FIGS. 12-16) and studies of the carotid artery(FIGS. 17-20). It will be appreciated by those of ordinary skill in theart that the description given herein with respect to these figures andexamples is for exemplary purposes only and is not intended in any wayto limit the scope of the invention. All questions regarding the scopeof the invention may be resolved by referring to the appended claims.

MR Angiography Using Steady-State Transport-Induced Adiabatic FastPassage Theory

Application of a constant, transverse RF field, B_(i), in the presenceof a magnetic field gradient, G_(i), results in adiabatic inversion ofspins which have a component of velocity in the direction of thegradient as illustrated in FIG. 1. In particular, FIG. 1 illustrates theeffect of a constant RF pulse on a spin moving through a gradient atconstant velocity. As shown, as a spin moves through the gradient, G_(i)(FIG. 1(a)), it undergoes a frequency sweep as its instantaneousfrequency changes linearly with time. As shown, the off-resonancecomponent Δω(k), where k = x, y or z, of the effective RF field, B_(i)(FIG. 1(b)), is swept from a large positive value to a large negativevalue (FIG. 1(c)). As will be shown below, if the rate of change of thefrequency is slow enough or if the strength of the RF signal is largeenough, the magnetization will follow the effective field and adiabaticinversion will occur. This general approach has previously been appliedto angiography and perfusion imaging by Dixon et al. and Williams et al.in the aforementioned patent application. As described therein, thestrength of the gradient, G_(i), and the velocity, v, of the spin definethe effective "sweep rate" for a moving spin in direct analogy tocontinuous-wave adiabatic fast passage. The present inventors have nowdiscovered that for adiabatic inversion to occur, the velocity shouldsatisfy the following equation: ##EQU1## where T₁ and T₂ are therelaxation times of the spin, γ is the gyromagnetic ratio, and B_(i) isthe strength of the off-resonance RF signal. Equation 1 illustrates thatif a spin moves too slowly, relaxation effects dominate and inversiondoes not occur, while conversely, if a spin moves too rapidly,incomplete inversion will occur because the sweep becomes non-adiabatic.The efficiency of an inversion pulse may thus be defined as:

    α=(M.sub.O -M.sub.Z)/2M.sub.O                        Equation 2

where M_(O) and M_(Z) are the z magnetizations before and after theinversion pulse, respectively. Complete inversion corresponds to α=1.

FIG. 2 illustrates the results of computer simulations of spins movingin the presence of such an inversion pulse. The computer simulation wasperformed using a spinor approximation of the Bloch equation. Asillustrated, the simulated inversion efficiency, α, is plotted versusthe strength, γB_(i), of the inversion pulse for three differentgradient strengths. For the purposes of calculation, the velocity of thespin was assumed to be 20 cm/sec. As expected, FIG. 2 illustrates thatthe inversion efficiency increases with increasing B_(i). FIG. 2 alsoshows that the inversion efficiency may be increased by decreasing thestrength of the inversion gradient.

The off-resonance RF pulse described above defines a plane of inversion.Any spin which moves through this plane will experience adiabaticinversion as long as its velocity satisfies the conditions ofEquation 1. The distance, d, between the inversion plane and isocenteris given by:

    d=ω.sub.i /(γG.sub.i),                         Equation 3

where ω_(i) is the frequency of the off-resonance pulse relative to thespectrometer frequency. The plane of inversion is perpendicular to theinversion gradient G_(i).

By preceding a standard imaging sequence with an off-resonance RF pulsein the presence of an inversion gradient in the slice-selectiondirection, the present inventors have discovered that steady-stateinversion of the arterial inflow to an organ may be achieved. Becausethe fraction of the arterial inflow which is inverted is directlyproportional to the duty cycle of the inversion pulse (i.e., the portionof TR including the inversion pulse), the inversion pulse should occupyas much of the interpulse interval as possible. An angiogram can then beformed by subtracting an image acquired with the inversion pulse presentfrom a control image acquired without arterial inversion.

As noted by Wolff et al. in an article entitled "Magnetization TransferContrast (MTC) and Tissue Water Proton Relaxation In Vivo", MagneticResonance in Medicine, Vol. 10, pp. 135-144 (1989), the presence ofmagnetization transfer in tissues complicates the proper choice of acontrol image. Indeed, because it is desired to use a single coil inaccordance with the invention, the off-resonance pulses used to effectinversion will partially saturate the broad component of the protonsignal, leading to magnetization transfer to spins in the image plane.This immediately implies that a control image may not be acquired simplyby reducing the amplitude of the inversion pulse to zero since themagnetization transfer effect would be completely unbalanced.Fortunately, as will be described below, approximate control for themagnetization transfer effect may be achieved by other means inaccordance with the invention.

For example, if M_(f) and M_(f) ^(O) are the equilibrium magnetizationsof unbound protons in the presence and absence, respectively, of anoff-resonance RF pulse of frequency ω_(i) and strength B_(i) and it isassumed that the inversion pulse is applied in the presence of agradient, G_(i), in the slice-selection direction, then the degree ofmagnetization transfer, μ, to unbound spins due to the off-resonancepulse may be defined as:

    μ= (M.sub.f O - M.sub.f) / (2M.sub.f O).                Equation 4

The dependence of μ on the frequency offset of the RF pulse, Δ=ω_(s)-ω_(i) (where ω_(s) is the resonance frequency of the spin in thepresence of the gradient) is given by:

    μ= μ(Δ) = K.sub.1 / (K.sub.2 +Δ.sup.2),  Equation 5

where K₁ and K₂ are constants which depend on the relaxationcharacteristics of the two spin populations as described by Grad et al.in an article entitled, "Nuclear Magnetic Cross-RelaxationSpectroscopy", Journal of Magnetic Resonance, Vol. 90, p. 1 (1990).These constants depend on the spin distribution and are thereforefunctions of position, that is, K₁ = K₁ (z) and K₂ = K₂ (Z). In thepresence of an inversion gradient, G_(i), in the z-direction, thefrequency offset becomes:

    Δ= γG.sub.i (z-d),                             Equation 6

where d is given by Equation 3 above. The degree of magnetizationtransfer in the inversion image, μ₁, as a function of position is thengiven by:

    μ.sub.i (Z) = K.sub.1 (Z)/[K.sub.2 (Z) + γ.sup.2 G.sub.i 2(z-d).sup.2 ].                                           Equation 7

If a "control" image is acquired with an inversion gradient of oppositesign or, equivalently, with the frequency offset of the inversion pulsereversed, the location of the plane of inversion changes sign to -d. Inthis case, the degree of magnetization transfer in the control image,μ_(c), as a function position is:

    μ.sub.c (Z) = K.sub.1 (Z)/[K.sub.2 (Z) + γ.sup.2 G.sub.i 2(z+d).sup.2 ].                                           Equation 8

Equations 7 and 8 imply that the difference in signal between controland inversion images is zero if and only if z=0, that is, for spins atisocenter. For spins which are not at isocenter in the direction of theinversion gradient (z is not equal to 0), the magnetization transfereffect is not properly controlled and incomplete background suppressionwill occur.

An example where the magnetization transfer effect is not properlycontrolled and hence incomplete background suppression will occur isillustrated in FIG. 3. In FIG. 3, it is assumed that the control imageis acquired by changing the sign of an inversion gradient in the slabselection direction. Due to dispersion of magnetization transfer in theslab selection direction, the same signal on inversion and controlimages is obtained only for spins at isocenter, i.e., z=0.

If, however, the spin distribution is symmetric across the slabthickness, then complete cancellation of the magnetization transfereffect will occur. In this case, K₁ (z) = K₁ (-z) and K₂ (z) = K₂ (-z),which implies that the difference signal, μ_(c) (z) - μ_(i) (z) is anodd function of position. Integration of this quantity across a slabcentered on isocenter yields zero residual signal. However, the presentinventors have found in practice that relatively good control for themagnetization transfer effect may be achieved in vivo by minimizing theslab thickness and the amplitude of the inversion pulse and bymaximizing the amplitude of the inversion gradient.

Thus, in accordance with the invention, the control image may begenerated either by changing the sign of the inversion gradient or bychanging the sign of the frequency offset, ω_(i), of the inversionpulse. Due to the presence of eddy currents which result from gradientswitching, it is preferable to change the sign of the frequency offset,for if one changes the sign of the inversion gradient, a differentpattern of eddy currents will exist in the inversion and control images,leading to incomplete background suppression.

Another effect which may cause decreased background suppression isactual deposition of the labeled blood in the capillary beds of, forexample, the brain parenchyma. This effect is the basis for anon-invasive perfusion imaging technique such as that described in theaforementioned related patent application. Fortunately, for the purposesof angiography, the magnitude of this effect is quite small, on theorder of only a few percent at 1.5 Tesla.

EXAMPLE 1

Experiments were conducted on a 1.5 Tesla MR imaging system (Signa, GEMedical Systems, Milwaukee) on human volunteers. The two-dimensionalaxial gradient-echo pulse sequence of FIG. 4 was used to acquireinversion and control images. As shown in FIGS. 4(a) and 4(b), anoff-resonance RF pulse with a constant magnitude, B_(i), and a constantgradient pulse, G_(i), preceded a standard imaging sequence. Thefrequency of the RF pulse, ω_(i), was chosen so that inversion alwaysoccurred through a plane 3 cm inferior to the imaging slab. Theinversion RF pulse is defined by its duration, T_(i), its strength,B_(i), and its frequency offset, ω_(i). It is preferably applied in thepresence of a field gradient G_(i) which is in the slab selectiondirection. As shown in FIG. 4(d), a homospoil y-gradient pulse is alsoapplied between the inversion and observation pulses to attenuateresidual transverse magnetization which develops as a result of theoff-resonance pulse.

FIG. 5 illustrates the locations of the image slab and inversion planeused to produce angiograms of the human brain in this example. In apreferred embodiment, the image slab is 30 mm thick and the inversionplane is located 3 cm inferior to the center of the slab, which is atisocenter. As shown, the inversion plane intersects the superior portionof the internal carotid artery as it enters the cranium. Atransmit/receive quadrature head coil was used to apply the RF inversionpulses and to acquire this image. In particular, the 30 mm thick imageslab was excited with a field-of-view of 22 cm, θ = 40°, TE = 5 msec andTR = 100 msec. A fractional echo was used to minimize artifacts arisingdue to dephasing during readout. A homospoil y-gradient pulse of 4 msecduration and amplitude of 1 gauss/cm is also applied between the end ofthe inversion pulse and the observation pulse as illustrated in FIG.4(d) to attenuate residual transverse magnetization created by theinversion pulse. For each slab, 256 phase encodings were performed andfour signal averages were acquired, yielding an acquisition time of 1minute and 43 seconds per image. The duration of the RF pulse was chosento be 80 msec, corresponding to a duty cycle (RF pulse duration / TR) of80%. Angiograms were then computed as the difference between magnitudeimages acquired with and without arterial inversion.

FIGS. 6(a) and 6(b) illustrate two-dimensional projective angiograms ofa patient produced using the technique of the invention. Theseangiograms were generated by using two different methods for control.Namely, FIG. 6(a) was generated from a control image acquired bychanging the sign of the inversion gradient, while the image of FIG.6(b) was generated from a control image acquired by changing the sign ofthe offset frequency of the inversion RF pulse. For both angiograms inFIG. 6, the inversion gradient was 0.5 gauss/cm, the amplitude of theinversion pulse was 80 Hz, and the offset inversion frequency, ω_(i),was -6400 Hz. The total acquisition time for each angiogram in FIG. 6was approximately 3.4 minutes.

Measurements of blood flow in the internal carotid artery (ICA) suggeststhat the average velocity is approximately 20 cm/sec as described byGullberg et al. Under these conditions, pulse simulations predict thatapproximately 80% inversion efficiency should occur in the ICA at 80 Hzas shown in FIG. 2. There is less background suppression in FIG. 6(a)due to the unbalanced eddy currents which result from the differentpattern of gradient switching. The images of FIG. 6 thus indicate thatthe changing of the sign of the offset frequency is a superior means ofcontrol.

FIGS. 6(a) and 6(b) also illustrate residual signal intensities near theventricles and the occipital cortex. These are regions where one wouldexpect to find a highly asymmetric spin distribution across the imageslab and hence poor control of the magnetization transfer effect.Nevertheless, the contrast-to-noise ratio within the proximal ICA is 41in FIG. 6(a) and 44 in FIG. 6(b).

FIG. 7 illustrates the effect of inversion gradient strength onangiographic contrast in accordance with the invention. In FIGS. 7(a)and 7(b), two angiograms of a patient were generated using two differentvalues for the inversion gradient. In FIG. 7(a), the inversion gradientG_(i) was 0.25 gauss/cm, while in FIG. 7(b) the inversion gradient G_(i)was 1.0 gauss/cm. For both angiograms the amplitude of the inversionpulse was 80 Hz and the total acquisition time for each angiogram was3.4 minutes. The plane of inversion was 3 cm inferior to the image slabin each case, and the inversion frequency, ω_(i), was -3200 Hz for FIG.7(a) and -1.2800 Hz for FIG. 7(b). As illustrated, better visualizationof the arteries is obtained at the higher gradient strength (FIG. 7(b))due to more efficient inversion. In fact, the pulse simulations of FIG.2 predict that the inversion efficiency for flow in the ICA under theseconditions should be approximately 0.5 for FIG. 7(a) and 1.0 for FIG.7(b). Moreover, the contrast-to-noise ratio within the proximal ICA wasmeasured to be 29 for FIG. 7(a) and 57 for FIG. 7(b). Also, at the lowervalue of the inversion gradient of FIG. 7(a), the magnetization transfereffects become more prominent, as indicated by poor backgroundsuppression near the occipital cortex and the ventricles. This isexpected due to greater dispersion of the effect across the slabthickness at the lower gradient strength.

FIGS. 8(a) and 8(b) illustrate the effect of inversion pulse amplitudeon angiographic contrast. Two angiograms of a patient were generatedusing two different values for the amplitude of the inversion pulse. Inparticular, the amplitude, B_(i), of the inversion pulse was 40 Hz forFIG. 8(a) and 120 Hz for FIG. 8(b). For both angiograms, the inversiongradient was 0.25 gauss/cm and the inversion frequency, ω_(i), was -3200Hz. The total acquisition time for each angiogram was 3.4 minutes. Asshown, there is better visualization of the arteries in FIG. 8(b) thanin FIG. 8(a) due to more efficient inversion of the arterial inflow.Indeed, pulse simulations predict that the inversion efficiency for flowin the ICA under these conditions should be approximately 0.5 for FIG.8(a) and 1.0 for FIG. 8(b). The contrast-to-noise ratio within theproximal ICA was also measured to be 30 for FIG. 8(a) and 64 for FIG.8(b). As expected, the magnetization transfer effects in the region ofthe occipital cortex and surrounding the ventricles are more prominentin the image generated with the higher RF (FIG. 8(b)). The averagespecific absorption rate for the image of FIG. 8(b) was determined to beapproximately 0.2 W/kg.

The use of a steady-state inversion technique such as that describedherein thus offers advantages over other techniques such as thatdisclosed by Dixon et al. which rely upon pulsed inversion.Specifically, one may shorten the TR interval in accordance with theinvention and use fast scan approaches to imaging without thecomplications of cardiac gating. In addition, the inversion pulse isinsensitive to variations in blood flow as long as the conditions ofEquation 1 are satisfied. Also, since the present invention does not useshaped pulses to label blood flow as in SIR techniques, no optimizedpulse shapes need to be computed and hence the demands on the RFamplifier hardware are less stringent. Moreover, while the motion of thespins during the inversion pulse is a source of error for techniqueswhich rely upon spatially selected pulses (SIR), the present inventionactually relies upon spin motion during the pulse for inversion tooccur. Yet another advantage of the invention is the fact that allmoving spins are inverted at the same location, which may be placedquite close to the imaging region. This reduces the loss of contrast dueto T_(i) relaxation of the labeled spins. Also, by using a single coilto apply the inversion and observation pulses, the blood supply may beinverted very close to the imaging slab so as to minimize transit timeeffects.

The amount of arterial inflow which is inverted, which should beproportional to angiographic contrast, is directly proportional to theduty cycle of the inversion pulse. For this reason, it is important touse as large a duty cycle as possible in accordance with the invention.In order to minimize the RF signal magnitude necessary for inversion itis also desired to reduce the amplitude of the inversion gradient.However, as the strength of the inversion gradient is decreased, failureto control for magnetization transfer effects leads to poorer backgroundsuppression in regions where high spin asymmetry across the imaging slabexists. Fortunately, the present inventors have found that, in practice,this effect is relatively small and does not significantly degrade thequality of the angiograms. Also, by using a single head coil to applythe pulses and by using small inversion gradients, it has been possibleto keep the average specific absorption rate well below FDA limits forhuman head exposures using the techniques of the invention.

Coronary Angiography Using Transport-Induced Adiabatic Fast PassageTheory

The MRA technique described above with respect to Example 1 has beenapplied to coronary vasculature by inverting the blood at an arbitrarylocation in the imaging volume proximal the heart. In particular, theprinciple of transport-induced adiabatic vast passage inversion has beenused to label blood flow in the coronary vessels. The basic technique isvery similar to that described above with respect to intracranialangiography except that, as will be described below, a "localizer" imageis first taken in order to determine the geometry of the patient'sheart.

For coronary angiography, as with intracranial angiography, a constantmagnitude, transverse RF field, B_(i), is applied in the presence of amagnetic field gradient, G_(i), so as to generate adiabatic inversion ofspins which have a component of velocity in the direction of thegradient. Thus, as described above with respect to Equation 1, if a spinmoves too slowly, relaxation effects dominate and inversion does notoccur, while conversely, if a spin moves to rapidly, incompleteinversion will occur because the sweep becomes non-adiabatic. However,computer simulations performed by the present inventors revealed thatthe RF amplitude necessary for 90% inversion of spins is relatively low.The inversion efficiency is a function of a dimensionless parameter, β,which is defined as:

    β=(γB.sub.i).sup.2 /(γG.sub.i v),         Equation 9

where γ is the gyromagnetic ratio and B_(i) is the amplitude of theinversion pulse. Computer simulations have also revealed thatapproximately 90% inversion is achieved when β is only two. As a resultof this, it is possible to use relatively low amplitude RF pulses toeffect inversion, thereby reducing the power deposition of theexperiment.

As described above with respect to Equation 3, the off-resonanceinversion pulse defines a plane of inversion. Any spin which movesthrough this plane will experience adiabatic inversion as long as itsvelocity satisfies the conditions of Equation 1. The distance betweenthe inversion plane and the isocenter is also given by Equation 3 above.The directional sensitivity of the inversion pulse is thus completelyunder operator control simply through choice of gradient direction.

FIGS. 9(a) - 9(c) illustrate how transport-induced adiabatic fastpassage pulses may be used to label blood flow in the coronary arteries.As shown in FIGS. 9(a) and 9(b), relative to the R wave of theelectrocardiogram (0 msec), the ventricular ejection period lasts fromapproximately 10 msec to 300 msec during systole. As shown in FIG. 9(c)coronary blood flow during this period is negligible due to myocardialcontraction. As will be appreciated by those skilled in the art, byapplying the inversion pulse during the period of ventricular ejection(T_(eject)), it is possible to effectively label all of the cardiacoutput. During diastole, labeled blood which is ejected from the leftventricle of the heart flows into the coronary arteries at the base ofthe aorta. Thus, by timing data acquisition to occur at the end ofdiastole, one may observe the labeled blood in the coronary arteries.Substraction of images acquired with and without the inversion pulseyields a coronary angiogram.

Such selective labeling of the ventricular ejection period isadvantageous for a number of reasons. First, it makes it possible tolabel the coronary blood flow with a shorter duration pulse, therebyreducing the duty cycle of the RF and the power deposition as comparedwith the embodiment described with respect to Example 1. In addition, itreduces errors which arise from the fact that the inversion pulse, as aresult of the gradient geometry, inevitably labels blood flow in theleft atrium. Labeled blood in the left atrium enters the left ventricleand causes a background signal to appear in the left ventricle in theangiogram. This "cavity effect" makes it harder to see the coronaryvessels in a projection angiogram. By shortening the duration of theinversion pulse and applying it only during ventricular ejection duringsystole, the "cavity effect" may be minimized. Moreover, the use of ashorter inversion pulse also reduces the amount of magnetizationtransfer to spins in the image slab. Magnetization transfer effects havebeen found by the present inventors to be an error in this form ofprojective angiography; therefore, by reducing the degree ofmagnetization transfer, use of a shorter inversion pulse improves thequality of the angiograms.

EXAMPLE 2

Experiments were conducted on a 1.5 Tesla MR imaging system (Signa, GEMedical Systems, Milwaukee) on human patients. The patients were placedin the bore of the magnet and a set of gated "localizer" images of theheart were acquired in approximately five minutes. These "localizer"images enabled the operator to determine from the parameters on thelocalizer image where to put the imaging plane and where to put theinversion plane. In particular, the "localizer" images allowed foraccurate localization of the left ventricle and aortic valve and allowedthe frequency offset of the inversion pulse to be determined usingEquation 3. A frequency was chosen such that blood flow in the superiorportion of the left ventricle would undergo inversion. In addition, thelocalizer images were used to determine the geometry of the image slab.In general, an image plane is desired which transects the aortic valveon its longitudinal axis so as to include the entire curve of the aorta,while an inversion plane is desired which is approximately onecentimeter inferior (intraventricle) with respect to the aortic valve.Also, the image slab preferably covers the entire myocardium.

The two-dimensional gated coronal gradient-echo pulse sequence of FIG.10 was then used to acquire inversion and control images and to formcoronary angiograms using transport induced adiabatic fast passage. Inparticular, a constant off-resonance RF pulse was applied in thepresence of a constant inversion gradient in the frequency encodingdirection as illustrated in FIGS. 10(b) and 10(e) prior to applicationof a standard two-dimensional gradient-echo imaging sequence asillustrated in FIGS. 10(b)-10(e). The inversion pulse was applied duringventricular ejection to label the ventricular outflow, and the image waslater acquired during diastole, by which time the labeled blood hadentered the coronary arteries. The frequency of the inversion pulse,ω_(i), was chosen so that inversion always occurred through a planeperpendicular to the z-axis and was positioned one centimeter inferior(intraventricle) to the aortic valve. The inversion pulse duration was240 msec (i.e., T_(i) = 240 msec) and it began 100 msec after the R wavetrigger as illustrated in FIGS. 10(a) and 10(b). The amplitude of theinversion pulse was 60 Hz and the amplitude of the inversion gradientwas 0.2 gauss/cm. Data acquisition then occurred 550 msec after the Rwave trigger as illustrated in FIGS. 10(b) - 10(e).

A transmit/receive quadrature body coil was used to apply the inversionRF pulse as well as the imaging pulse, and a 40 mm thick slab over themyocardium was excited with a field-of-view of 24 cm, θ =90° and TE = 5msec. A fractional echo was used to minimize artifacts arising due todephasing during readout. A homospoil y-gradient pulse of 4 msecduration and amplitude of 1 gauss/cm was also applied between the end ofthe inversion pulse and the observation pulse to attenuate residualtransverse magnetization created by the inversion pulse in the preferredembodiment (FIG. 10(c)). Four signal averages were acquired forinversion and control images yielding a total acquisition time ofapproximately 14 minutes. The average specific absorption rate of thisexperiment was 0.3 W/kg, which is below the FDA limit for human wholebody exposures. Angiograms were then computed as the difference betweenmagnitude images acquired with the amplitude of the inversion pulse setto 60 Hz and 0 Hz, respectively.

FIG. 11 illustrates an angiogram produced in accordance with thetechnique of the invention. The bright horizontal band in FIG. 11indicates the geometry of the inversion plane. The left anteriordescending (LAD) branch of the left main coronary artery is alsolabeled. Blood flow in the proximal left coronary artery is clearlylabeled by the inversion pulse, while blood within the left ventricle isalso labeled due to residual cavity effects. Unfortunately, in FIG. 11this makes it difficult to observe the course of the left coronaryartery as it descends in front of the left ventricle.

Example 2 demonstrates that coronary blood flow may be labeled throughuse of transport-induced adiabatic fast passage inversion pulses. Asnoted above with respect to Example 1, use of this technique in asteady-state fashion in a human brain revealed the presence ofmagnetization transfer effects which lead to incomplete backgroundsuppression. However, the present inventors have found that theseeffects are negligible in the coronary circulation, which is most likelydue to the fact that the duty cycle of the inversion pulse was onlyabout 20% in Example 2, in contrast to duty cycles of 80% or more whichwere used in the cerebral applications described above with respect toExample 1.

Improvements in the quality of the angiograms in Example 2 may beachieved by pursuing alternate geometries to better visualize thecoronary vessels. In addition, more careful timing of the inversionpulse should further minimize cavity effects. Specifically, it may bedesirable to only label the latter half of the systolic ejection sincemost of the blood which enters the coronary circulation is ejectedduring this time period. The use of higher resolution flow compensationgradient waveforms, projection imaging methods, single-shot imaging(echoplanar imaging) and/or respiratory compensation should also improvethe quality of the angiograms of Example 2 in accordance with theinvention.

Quantitative MR Flow Measurement Using Pulsed Adiabatic Inversion Theory

By using the MR angiography techniques described above with respect toExamples 1 and 2 in conjunction with the tissue perfusion techniquedescribed in the aforementioned U.S. patent application and a techniquefor MR blood flow measurement, a comprehensive MR imaging system forimaging the heart, the brain, the kidney, the liver and other majortissues and organs may be accomplished nonintrusively using magneticresonance. In order to fulfill the requirements of such a comprehensiveimaging system, the present inventors have further developed a techniquefor measuring flow velocity using constant amplitude RF pulses in thepresence of a magnetic gradient to adiabatically invert spins which movein the direction of the gradient as in the techniques described aboveand by Williams et al. in the aforementioned U.S. patent application andpaper.

The theory of MR flow measurement in accordance with the invention issimilar to that given above for MR angiography. In particular, asdescribed above with respect to FIG. 1, a spin which moves through alinear magnetic field gradient at constant velocity in the presence ofan off-resonance RF pulse undergoes adiabatic fast passage by virtue ofits motion. FIG. 12 illustrates the effect of a constant RF pulse on aspin moving through a gradient at constant velocity. As shown in FIG.12, as the spin moves through the gradient, G, it undergoes a frequencysweep which has a rate determined by the velocity of the spin and thestrength of the gradient. If the sweep rate is slow enough or if thestrength of the RF pulse is large enough, the magnetization will followthe effective field, B_(eff), and adiabatic inversion will occur. Inshort, adiabatic inversion of the liquid will occur if its velocitysatisfies Equation 1. As noted above, this approach to spin labelingdiffers fundamentally from approaches using spatially selective pulsesin that it relies upon motion of the spins during the pulse to effectinversion.

The inversion gradient defines an "inversion plane" perpendicular to thegradient whose location d is given above by Equation 3. However, in thisembodiment, application of intermittent pulses of RF in the presence ofa gradient results in "bands" of inversion in the fluid which propagatedownstream from the site of inversion. The length, Δ, of an inversionband is given by:

    Δ = v · T.sub.i,                            Equation 10

where Ti is the duration of the inversion pulse. As will be describedbelow, in a pulsed steady-state imaging experiment a characteristicbanding pattern is established in the fluid.

FIG. 13(a) illustrates a typical banding pattern established in aflowing liquid by a train of RF pulses as illustrated in FIG. 13(b).Each band corresponds to a specific timing interval in the pulsesequence of FIG. 13(b), where T_(i) is the duration of the inversionpulse, T_(ed) is the time interval between the end of the inversionpulse and the center of the echo, and TR is the repetition time.Measurement of the length of the bands allows calculation of the averagevelocity over a known time interval from Equation 10. By gating such apulse sequence to the cardiac cycle, it is possible to make makenumerous quantitative in vivo measurements of flow velocities in asingle cardiac cycle and to average the measured velocities over thecycle.

EXAMPLE 3

An experiment was conducted on a simple phantom and on humans using a1.5 Tesla GE Signa imaging system.

As illustrated in FIG. 14, an ungated two-dimensional coronal gradientecho pulse sequence was used to implement the technique of the inventionto study flow in a simple phantom. In FIG. 14(a), a constant magnitude,off-resonance inversion RF pulse defined by its duration, T_(i), itsstrength, B_(i), and its frequency offset, ω_(i), was applied in thepresence of a field gradient G_(i) which in the sequence of FIG. 14 isin the z-direction (FIG. 14(d)). A homospoil gradient pulse in thex-direction having an amplitude of 1 gauss/cm and a width of 4 msec wasalso applied (FIG. 14(b)) after the inversion pulse to attenuateresidual transverse magnetization as a result of the off-resonancepulse. A standard imaging sequence was then applied.

A transmit/receive quadrature birdcage head coil was used for applyingthe inversion pulse and acquiring the image. Magnitude, coronal imageswere acquired with θ = 30°, TE = 6 msec, TR = 80 msec, a field-of-view(FOV) of 16 cm, and a slice thickness of 3 mm. A fractional echo wasalso used to suppress flow artifacts arising from dephasing duringreadout. The phantom consisted of 0.5 inch diameter Tygon tubing throughwhich tap water was driven in the +z direction by a varistaltic pump.The slice location was chosen to be at the center of the tube. The flowrate was determined to be 520 ml/min by collecting a volume over a knowntime interval.

FIGS. 15(a) - 15(c) illustrate typical banding patterns produced by thepulse sequence of FIG. 14 for three different strengths of the inversionpulse: B_(i) = 20, 40, and 80 Hz, respectively. In particular, FIGS.15(a) - 15(c) illustrate magnitude gradient echo images of water flowingthrough 1/2 inch diameter tubing at a flow rate of 520 ml/min in the +zdirection, corresponding to a peak velocity of about 14 cm/sec. A 40msec inversion pulse was applied 2554 Hz below resonance in the presenceof a z-gradient of 0.1 gauss/cm. The inversion pulse was thuscharacterized by G_(i) = 0.1 gauss/cm, ω_(i) = -3000 Hz and T_(i) = 40msec. The echo delay, T_(ed), was chosen to be 10 msec. As shown inFIGS. 15(a) - 15(c), as the RF signal intensity increases, the degree ofinversion increases, reflected by the development of bright inversionbands (at the arrows) at the higher magnitude images of FIGS. 15(b) and15(c). Pulse simulations performed by the present inventors predictedthat the inversion efficiency, α, is as defined in Equation 2, whereM_(O) and M_(Z) were at the z-magnetizations before and after passagethrough resonance and should be 0.50, 0.90 and 0.95 for inversionstrengths of 20, 40, and 80 Hz, respectively. Values for α were measuredfrom the images of FIGS. 15(a) - 15(c) by computing the ratio of theimage intensity in the center of the first inversion band to the imageintensity just upstream from the inversion plane. The values obtainedwere 0.38, 0.76 and 0.86 for FIGS. 15(a), 15(b), and 15(c) respectively.As shown, the intensity of the inversion bands decreases due to T1relaxation as the flow propagates downstream, with the inversion bandsof FIG. 15(c) lasting longer than those of FIG. 15(b) due to morecomplete inversion. This effect may, in part, explain why the measuredvalues for α are somewhat lower than the predicted values.

In FIG. 16, the measured (solid) and predicted (dashed) values for theflow velocity calculated from FIG. 15(c) are plotted versus distanceacross the lumen for a steady flow in a 1/2 inch diameter tube phantom.The flow velocity was determined by measuring the thickness of the firstinversion band and dividing by the time interval T_(i). Also plotted isthe laminar flow profile which corresponds to the measured volumetricflow rate and tube diameter. The Reynold's number defined by the peakvelocity and tube diameter in the illustrated example is approximately1700, indicating that laminar flow should exist. Excellent agreementbetween predicted and measured values is obtained except in the regionwhere the velocity is low, such as near the vessel wall.

The images of FIG. 15 illustrate the usefulness of the technique of theinvention for observing the dynamic aspects of flow. As shown, as thespin boluses propagate downstream from the inversion plane, contrastbetween bands is lost. This loss of contrast results from the effects ofT1 relaxation of the spins as they travel downstream and mixing whichoccurs as a result of diffusion. This spins immediately next to thevessel wall, which are moving too slowly to undergo inversion, aresimply saturated. As they flow downstream, they recover via T1relaxation over a shorter distance than the more rapidly flowing spins,and hence are relatively bright. The fact that they retain higher signaldown the length of the tube implies that little mixing between layers isoccurring, as one would expect for laminar flow.

Human Studies

Cardiac gated flow velocity waveforms were acquired from the internalcarotid artery (ICA) of patients using the modified Cine pulse sequenceof FIG. 17. FIG. 17 illustrates a two-dimensional gated pulse sequenceused to make in vivo flow measurements, whereby after the R wavetrigger, a series of images are acquired, each corresponding to adifferent time point of the cardiac cycle. As shown in FIG. 17(e), therepetition time, TR, was chosen to be as small as possible in order tomaximize the number of observed cardiac phases. Magnitude coronal imageswere formed with θ =30°, TE = 5 msec (a fractional echo was used), a 16cm FOV, and a slice thickness of 12 mm. A constant off-resonance RFpulse (FIG. 17(b)) in the presence of a z-gradient (FIG. 17(e)) precededeach observation pulse.

Flow velocities were then calculated by measuring the distance betweenthe plane of inversion and the trailing edge of the first inversion bandand dividing by the time between the end of the inversion pulse and thecenter of the echo, the echo delay T_(ed) as illustrated in FIG. 14.These measurements yielded average values for the velocity of the spinsduring the time interval T_(ed). However, those skilled in the art willappreciate that the average velocity could also be calculated bymeasuring the length of the first inversion band and dividing by thetime interval T_(i). Each gated image therefore contains informationabout the velocity during more than one part of the cardiac cycle.

FIGS. 18(a) and 18(b) illustrate systolic and diastolic images,respectively, from a set of 15 cardiac phases acquired from the neck ofa patient. These gated magnitude gradient-echo images of the internalcarotid artery of the patient were obtained using the pulse sequence ofFIG. 17. The image plane was chosen so that flow in the internal carotidarteries (ICA) could be measured. Due to the fact that the slicethickness was greater than the diameter of the carotid artery, theobserved banding pattern represented the integral of the spindistribution across the artery. The inversion pulse was thuscharacterized by G_(i) = 0.2 gauss/cm, ω_(i) = -2554 Hz and T2i = 20msec. The echo delay, T_(ed), was chosen to be 10 msec, while thestrength of the inversion pulse was 80 Hz. The position of the labelingplane is indicated by the horizontal saturation band in the images. Asillustrated, there is significantly greater superior displacement of thefirst labeled band in the systolic image (FIG. 18(a)) than in thediastolic image (FIG. 18(b)). Blood flow in the internal jugular veinsis also tagged by the inversion pulse, as indicated by the inferiordisplacement of a saturation band lateral to the ICA.

Values for peak velocity in the left and right ICA were determined bymeasuring the maximal length of the first inflow band for each image anddividing by the echo delay, T_(ed). This yielded a total of 15measurements of flow velocity. The peak velocity measurements for theright ICA are plotted versus trigger delay in FIG. 19, which illustratesflow velocity as a function of trigger delay in the right internalcarotid artery of a patient obtained by measuring the thickness of thefirst inflow band in a series of images gated to the cardiac cycle. Themeasurements for the left ICA are nearly identical. As shown, the peakblood flow is approximately 70 cm/sec. As shown, the velocity in the ICAnever becomes 0, and instead approaches a minimal value of approximately10 cm/sec. A minimum in the velocity occurs approximately 350 msec afterthe R wave (after systole). These results generally agree with thefindings of Firmin et al. in an article entitled "Echo-PlanarHigh-Resolution Flow Velocity Mapping", Magnetic Resonance in Medicine,Vol. 12, pp. 316-327 (1989).

The flow measurement technique of the invention thus provides a simpleand robust way to measure fluid flow. It differs fundamentally fromother techniques in that motion of the spins during application of theRF pulses is necessary for labeling to occur and in that the width ofthe inversion bands is directly proportional to the time duration of theinversion pulse. This is in direct contrast to prior art methods whichuse spatially selective tagging pulses where the thickness of the bandis inversely proportional to the time duration of the tagging pulse.

The aforementioned technique of the invention also uses very simple,constant amplitude pulses. Precise band definitions are made possibledue to the short frequency interval over which adiabatic inversiontypically occurs. FIG. 20 illustrates pulse simulations of the effect ofincreasing gradient strength on adiabatic flow inversion. In FIG. 20,values of the z-magnetization of a spin moving through a gradient in thepresence of a constant RF pulse is plotted as a function of distancefrom resonance for three different values for the field gradient, G_(i): 0.1 (solid), 0.2 (dotted) and 0.4 gauss/cm (dashed). For thesesimulations, the spin velocity was 10 cm/sec and the amplitude of the RFpulse was 60 Hz. As illustrated, at the highest values of the inversiongradient, the distance over which 90% inversion occurs is smallest, lessthan 1 mm. For lower values of the gradient, the transition becomes lesssharp, implying that the banding pattern would be less distinct. FIG. 20also illustrates, however, that inversion efficiency decreases at thehigher gradient strength. Thus, as will be recognized by those skilledin the art, there is generally a trade-off between inversion efficiencyand band definition.

An intrinsic limitation of this technique is its inability to measureslow flow. This results from the fact that relaxation effects becomesignificant at the lower sweep rates defined by low velocities. Forexample, if it is assumed that T₂ = 300 msec, G_(i) = 0.1 gauss/cm andB_(i) = 80 Hz, then Equation 1 predicts that inversion will fall off forvelocities lower than about 1 cm/sec.

RF exposure of the patient may also be minimized in accordance with thetechnique of the invention by choosing as small a gradient as possiblefor inversion, while at the same time maintaining reasonable definitionbetween bands. For the gated pulse sequence used in the experiments ofExample 3, the average specific absorption rate was approximately 0.05W/kg, while the peak specific absorption rate was approximately 2.5W/kg.

The flow measurement technique of the invention thus allows for directobservation of the propagation of many boluses following inversion. Theability to observe the propagation of a larger number of boluses allowsone to observe the dynamic effects of turbulence and complex flow. Otherbenefits of the invention will be apparent to those with ordinary skillin the art.

Although exemplary embodiments of the invention have been described indetail above, those skilled in the art will readily appreciate that manyadditional modifications are possible in the exemplary embodimentswithout materially departing from the novel teachings and advantages ofthe invention. For example, the sequential phase encoding technique ofthe invention may be used for MR angiography and flow measurement ofmany other organs, tissues and arteries of the body, including, forexample, the kidney, the liver, the lungs, the major arteries andmuscles, and the like. In addition, the technique of the invention maybe modified by using flow compensated gradient waveforms duringacquisition of the observed signal, thereby preserving the signals fromthe arteries. Those skilled in the art will also appreciate that the MRangiography and blood flow measurement techniques herein described maybe combined with the tissue perfusion techniques described in theaforementioned patent application to Williams et al. to provide MRpackages which may be used in comprehensive heart, brain or other organexaminations. Accordingly, all such modifications are intended to beincluded within the scope of this invention as defined in the followingclaims.

We claim:
 1. A method of generating an angiographic image of a bodyportion of a patient using magnetic resonance, comprising the stepsof:applying to the body portion an external substantially uniformmagnetic field so as to align predetermined nuclei of said body portionwith a constant magnetic field gradient of said magnetic field, saidpredetermined nuclei having a particular resonance frequency; definingan inversion plane in said body portion which intersects a blood supplyin a direction transverse to said magnetic field gradient by applyingvia an RF coil disposed adjacent to said body portion a substantiallyconstant amplitude tagging radio-frequency (RF) pulse having a frequencydifferent from said particular resonance frequency, said tagging RFpulse being applied in the presence of said constant magnetic fieldgradient so as to invert spins of those predetermined nuclei which havea component of velocity in a direction of said magnetic field gradientas said predetermined nuclei pass through said inversion plane; phaseencoding said predetermined nuclei as they pass through said inversionplane during said tagging RF pulse so as to establish a steady stateinversion condition for said predetermined nuclei; applying via said RFcoil an observation RF pulse for acquiring a tagged image of said bodyportion in a region near said inversion plane; acquiring an untaggedimage of said region; and effectively subtracting said tagged image ofsaid region from said untagged image of said region to obtain anangiogram of said body portion in said region.
 2. A method as in claim1, wherein said tagging RF pulse is applied during said defining stepfor at least 80% of an RF pulse period defined as the total duration ofsaid tagging pulse and said observation RF pulse.
 3. A method as inclaim 2, wherein the RF pulse period is no more than 100 msec and isasynchronous with a cardiac cycle of the patient.
 4. A method as inclaim 1, wherein said step of acquiring said untagged image of saidregion comprises the step of changing a sign of a frequency offset ofsaid tagging RF pulse prior to application of an imaging sequence.
 5. Amethod as in claim 1, wherein said defining step includes the step ofinverting the spins of said predetermined nuclei which have a componentof velocity, v, in the direction of said magnetic field gradient whichsatisfy the relationship: ##EQU2## where G_(i) is the magnitude of saidmagnetic field gradient, B_(i) is the magnitude of said tagging RFpulse, T₁ and T₂ are relaxation times of the spins of said predeterminednuclei, and γ is the gyromagnetic ratio.
 6. A method as in claim 1,comprising the further step of applying a homospoil gradient pulsebetween said tagging and observation RF pulses for attenuatingtransverse magnetization which developes as a result of said tagging RFpulse.
 7. A method as in claim 1, wherein said inversion plane isdefined in said defining step so as to intersect a superior portion ofthe internal carotid artery as it enters the cranium of the patient. 8.A method as in claim 1, wherein said inversion plane is defined in saiddefining step so as to intersect the heart at a plane approximately 1 cmwithin the left ventricle with respect to the aortic valve of thepatient.
 9. A method as in claim 8, wherein said tagging RF pulse isapplied during a predetermined portion of the duration of ventricularejection and said observation RF pulse is applied at the end of diastoleso that phase encoded nuclei in the coronary arteries of the heart areacquired in said tagged image.
 10. A method as in claim 8, wherein saidinversion plane defining step comprises the further steps of obtainingat least one gated localizer image of the heart, determining thelocation of the left ventricle and aortic valve of the heart from thelocalizer image, determining a frequency offset of said tagging pulsefrom the localizer image, and determining from the localizer image adesirable location for an imaging plane for acquiring said tagged imageprior to application of said tagging RF pulse.